In at least one known CT system configuration, an x-ray source projects a fan-shaped beam which is collimated to lie within an X-Y plane of a Cartesian coordinate system and generally referred to as the "imaging plane". The x-ray beam passes through the object being imaged, such as a patient. The beam, after being attenuated by the object, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is dependent upon the attenuation of the x-ray beam by the object.
Each detector element of the array produces a separate electrical signal that is a measurement of the beam attenuation at the detector location. Particularly, each x-ray detector element typically includes a collimator for collimating x-ray beams received at the detector cell, and a scintillator is located adjacent the collimator. The scintillator includes a plurality of scintillating elements, and adjacent scintillators are separated by a non-scintillating gap. Photodiodes are positioned adjacent the scintillator elements and generate electrical signals representative of the light output by the scintillator elements. The attenuation measurements from all the detector cells are acquired separately to produce a transmission profile.
In known third generation CT systems, the x-ray source and the detector array are rotated with a gantry within the imaging plane and around the object to be imaged so that the angle at which the x-ray beam intersects the object constantly changes. A group of x-ray attenuation measurements, i.e., projection data, from the detector array at one gantry angle is referred to as a "view". A "scan" of the object comprises a set of views made at different gantry angles during one revolution of the x-ray source and detector. In an axial scan, the projection data is processed to construct an image that corresponds to a two dimensional slice taken through the object. One method for reconstructing an image from a set of projection data is referred to in the art as the filtered back projection technique. This process converts that attenuation measurements from a scan into integers called "CT numbers" or "Hounsfield units", which are used to control the brightness of a corresponding pixel on a cathode ray tube display.
To reduce the total scan time, a "helical" scan may be performed. To perform a "helical" scan, the patient is moved while the data for the prescribed number of slices is acquired. Such a system generates a single helix from a one fan beam helical scan. The helix mapped out by the fan beam yields projection data from which images in each prescribed slice may be reconstructed.
Multislice CT systems are used to obtain data for an increased number of slices during a scan. Known multislice systems typically include detectors generally known as 2-D detectors. With such 2-D detectors, a plurality of detector cells form separate columns, or channels, and the columns are arranged in rows. Each row of detectors forms a separate slice. For example, a two slice detector has at least two rows of detector cells, and a four slice detector has at least four rows of detector cells. During a multislice scan, multiple rows of detector cells are simultaneously impinged by the x-ray beam, and therefore data for several slices is obtained.
In a multislice detector, each cell is subjected to X-rays from a range of angles depending on its Z-axis location. In one configuration for a helical scan, the angular range is about .+-.1 degree at the extreme Z axis edges of the detector. Because the detector is made up of scintillating segments separated by small non-scintillating gaps, the signal will be at a minimum when the X-ray beam is generally perpendicular to the scintillators. The signal increases as the angle of the x-ray beam increases from perpendicular because the perpendicular X-ray beams have the lowest geometric collection efficiency. Such low geometric efficiency results since angled X-rays are presented with a smaller effective non-scintillating gap than perpendicular X-rays.
Further, since each channel (or sets of channels built as modules), do not have identical gap configurations or Z-axis positions within the detector, there is a phase difference between the minimum gain points. Additionally, the focal spot typically moves in the Z-axis up to 1.0 mm due to the thermal expansion and centrifugal forces interacting with gravity. This position change creates a change in incident angle by about 0.06 degrees. Due to the phase differences between the minimum gain points, differential channel gain variations of 0.2% or more can occur over the range of focal spot positions.
Third generation CT scanners may produce ring, band and center spot artifacts when differential gain errors exceed 0.02%. Differential gain values are calibrated and then corrected during image reconstruction. However, the variation in incident angle changes the differential gain during scanner operation and hence cannot be easily corrected with software algorithms.
It would be desirable to provide a scintillator construction that has increased geometric efficiency compared to known scintillator constructions. It also would be desirable to provide such a scintillator construction which does not increase the scintillator fabrication costs nor reduce dose efficiency of the system.